Medical diagnostic imaging began with the discovery of x rays by W. C. Roentgen in 1895 and today includes radiography, nuclear medicine imaging, ultrasound imaging, computed tomographic imaging, and magnetic resonance imaging. In general the goal of each type of medical imaging is to provide a spatial mapping of a parameter, feature, or process within a patient.
In radiology and computed tomography, a source of x rays is beamed through the patient onto a suitable detector such as a film or a plate. The detector measures the intensity distribution of the incident beam of x rays and provides an image representing the attenuation of the radiation resulting from the absorption and scattering within the patient's body.
Nuclear medicine involves injection of a radiopharmaceutical into a patient and measurement of the intensity distribution of gamma radiation emitted from the patient's body. Radiopharmaceuticals are formed by attaching a radioactive tracer to a pharmaceutical that is known to preferentially accumulate in the organ of interest. Thus, the radiation pattern is a measure of blood flow, metabolism, or receptor density within the organ of interest and provides information about the function of the organ. Either a single projection image of the radiation pattern may be taken (planar imaging) or many projection images may be acquired from different directions and used to compute the three dimensional emission distribution (single photon emission computed tomography, or SPECT). Radiation-imaging systems used in nuclear medicine are often referred to as "gamma" cameras.
Pioneer nuclear medicine imaging systems used scanning methods to generate images. Such pioneer systems generally used a scintillation-type gamma-ray detector equipped with a focusing collimator which moved continuously in selected coordinate directions, i.e., in a series of parallel sweeps, to scan regions of interest. A disadvantage of these early imaging systems was the lengthy exposure times that were required to derive an image of the system or organ under test. In addition, dynamic studies of such organs were often difficult to obtain.
Another type of prior art radiation detection system utilizes an "Anger" type gamma scintillation camera (named after its inventor H. O. Anger, see "A New Instrument for Mapping Gamma Ray Emitters," Biology and Medicine Quarterly Report, U.C.R.L.-3653, 1957), for determining the radiation pattern emitted from a patient's body. These nuclear medicine imagers use large sodium iodide scintillating crystals in conjunction with a bank of photomultiplier tubes (PMTs). A collimating aperture in front of the scintillation crystal focuses the gamma rays on the crystal, and gamma rays from a radiopharmaceutical injected into the patient produce light flashes (scintillations) in the crystal which are converted into electrical signals by the PMTs. High density shielding material, typically lead, is used to cover the sides and back of the radiation detection assembly to prevent radiation from entering the detector by any path other than through the collimator. A computer locates each flash from the relative magnitudes of the PMT signals. Crystals are typically 200 to 400 square inches in area.
Limitations in the Anger camera stem from the process of converting scintillations into electrical signals. Sources of distortion include: 1) variation of the acceptance field-of-view angle of the PM tubes with distance from the scintillation event, 2) refraction and light guiding due to index of refraction mismatches, 3) unavoidable dead regions between PMTs, 4) higher effective density (hence, heavier weighting) of distant PMTs, 5) non-uniform spatial response of individual PMTs, 6) variation in response from one PMT to another, 7) temporal variation of PMT response, and 8) an unavoidable dead margin several centimeters wide around the perimeter related to the inability of determining positions outside the middle of the outer PMTs. Other errors stem from instabilities in the PMTs and the fragility and hygroscopic nature of the scintillation crystal.
Disadvantageously, because of the large size of the detection assembly that results from the combination of scintillator, light pipe, and photomultiplier tubes, the lead shielding dramatically increases the weight and cost of Anger cameras. Furthermore, the non-sensitive (dead) margin around the perimeter of the Anger camera makes it difficult to adequately image small organs and some body parts (the breast, for example). In addition, the large size of the Anger camera and its weight prevent it from being used effectively in locations such as in operating rooms, intensive care units, or at the patient's bedside.
Inherent to the Anger camera design, the scintillator detection element is formed in a plane. There could be significant advantage for some applications of forming the detection elements in a shape that conforms more closely to that of an object to be imaged.
Semiconductor detector-array imagers have been proposed for solving problems with Anger cameras, e.g., see U.S. Pat. No. 4,292,645; U.S. Pat. No. 5,132,542; IEEE Transactions on Nuclear Science, Vol. NS-27, No. 3, June 1980, "Semiconductor Gamma Cameras in Nuclear Medicine;" and IEEE Transactions on Nuclear Science, Vol. NS-25, No. 1, February 1978, "Two-Detector, 512-Element, High Purity Germanium Camera Prototype." It has long been recognized that semiconductor detector arrays are potentially attractive for nuclear medicine imaging because of their very small size and weight, excellent spatial resolution, direct conversion of gamma photons into electrical signals, capability of on-board signal processing, high stability, and reliability. Using this technique, gamma-ray radiation absorbed in a semiconductor detector produces holes and electrons within the detector material which, due to the influence of a bias voltage, separate and move toward opposite surfaces of the semiconductor material in accordance with their respective electrical charge polarities. The electron and hole currents are then amplified and conditioned by electronic circuitry to produce electrical signals which are processed to indicate the location and intensity of the corresponding incident gamma-ray radiation.
Prototype semiconductor detector-array cameras embodying these principles have been developed with varying degrees of success. For example, attempts at using two-dimensional detector arrays of cryogenically-cooled-germanium detectors and room-temperature HgI.sub.2 detectors have generally been limited to the scientific laboratory due to the problems associated with cryogenic cooling and practical difficulties with HgI.sub.2 technology. An early feasibility study of an imaging system based on a rotating linear array of cadmium telluride (CdTe) detectors has similarly not proven to be a satisfactory solution and has apparently been abandoned.
One example of a prior art semiconductor gamma camera is described in U.S. Pat. No. 4,292,645, to Schlosser, et al. Schlosser teaches an improved technique for providing the necessary electrical contact to doped regions of a semiconductor gamma detector principally comprised of germanium. A layer of resistive material makes contact with conductive strips on the detector surface, and two readout contacts at the sides of the resistive layer, parallel to the strips and connected to two amplifiers, allow identification of the strip where a gamma ray is absorbed. The opposite side of the detector is arranged the same except that the strips are orthogonal to those on the top. The spatial position of an event is the intersection of the identified orthogonal strips. Two amplifiers for the top surface and two amplifiers for the bottom surface handle all events in the entire imager. Though this keeps the electronic component count small, it is a disadvantage to use the entire crystal for detection of each gamma ray. As a result of this, the resolution gets worse and the achievable count rate decreases as the size of the detector is increased.
Another example of a prior-art gamma-ray-imaging system using a semiconductor detector array is described in Materials Research Society Symposium Proceedings, Vol. 302 (Materials Research Society, Pittsburgh, 1993), pp. 43-54, "Multi-Element Mercury Iodide Detector Systems for X-Ray and Gamma-Ray Imaging," by Bradley E. Patt. Patt teaches the use of orthogonal strips on opposite sides of the semiconductor crystal to define the semiconductor detector array pixels, with one amplifier being used for each strip. The coincidence of signals from orthogonal strips is used to define the position at which a gamma ray is absorbed within the crystal. Disadvantageously, as the area of the detector gets larger and the length of the strips increases, the capacitance associated with the strip and the leakage current in the strip from the detector increase. Both capacitance and leakage current reduce the pulse energy resolution which degrades the imager performance.
The prior art lacks a semiconductor detector array that is large enough to satisfy nuclear medicine applications or that operates at room temperature. Therefore, there is a need for a detector which overcomes the disadvantages of the Anger camera, has an active area appropriate for medical imaging application, has negligible dead region around the perimeter, and operates at room temperature. There is a need for a cost-effective means of manufacturing such detectors for nuclear medicine and other applications.
A semiconductor detector array may be realized by combining together many individual detector elements. However, when the individual detector elements are made sufficiently small to meet spatial resolution requirements, the number of amplifiers needed to amplify the signals becomes very large. For infrared and low-energy x-ray applications, prior art focal-plane arrays and silicon-strip detectors combine amplifiers for each element and a multiplexer that provides a single output for the large number of inputs (see Nuclear Instruments and Methods in Physics Research, Vol. 226, 1984, pp. 200-203; and IEEE transactions on Nuclear Science, Vol. NS-32, No. 1, February 1985, p 417). These prior art readout circuits are not adequate for handling signals produced by gamma-ray detectors such as CZT detector arrays required for nuclear medicine imaging.
In addition, because of variations in response between individual detector elements and between individual amplifiers, a need exists for a method to normalize the gain and efficiency of each detection element and its associated amplifier.
The present invention provides such a semiconductor gamma-ray camera and imaging system wherein both planar images and SPECT images may be obtained. The imaging system includes a detector for sensing radiation emitted from a subject under test, electronics for conditioning and processing the detected radiation signals, a computer for controlling the detection process and of forming and displaying images based upon the signals generated by the detectors, and output devices for displaying the images and providing data to a user.